This invention relates to x-ray or gamma ray detector arrays which are useful in computerized tomography scanners or related industrial applications. More specifically, the invention relates to a scintillation detector array in which the collimator/channel separator plates are arranged in a zig-zag configuration.
Computerized tomography (CT) scanners are medical diagnostic instruments in which the subject is exposed to a relatively planar beam or beams of x-ray radiation, the intensity of which varies in direct relationship to the energy absorption along a plurality of subject body paths. By measuring the x-ray intensity (i.e., the x-ray absorption) along these paths from a plurality of different angles or views, an x-ray absorption coefficient can be computed for various areas in any plane of the body through which the radiation passes. These areas typically correspond to an approximately square portion having dimensions of approximately 1 mm.times.1 mm. The absorption coefficients are used to produce a display of, for example, the bodily organs intersected by the x-ray beam.
An integral and important part of the scanner is the x-ray detector which receives the x-ray radiation which has been modulated by passage through the particular body under study. Generally the x-ray detector contains a scintillation material which, when excited by the impinging x-ray radiation, emits optical wavelength radiation. In typical CT or industrial applications, the optical output from the scintillator material is made to impinge upon photoelectrically responsive materials in order to produce an electrical output signal. The amplitude of this signal is in direct relation to the intensity of the impinging x-ray radiation. The electrical signals are converted to a digital form to be processed by digital computer means which generates the absorption coefficients in a form suitable for display on a cathode ray tube screen or other permanent media.
In the past, a parallel plate collimator/channel separator construction has been employed in scintillation detector arrays for use in CT. Two examples of such arrays are described, respectively, in U.S. Pat. No. 4,187,427 issued on Feb. 5, 1980 to D. A. Cusano (one of the inventors herein) and in U.S. Pat. No. 4,262,202 issued Apr. 14, 1981 to the same inventors as herein. The above-identified patents are assigned to the same assignee as the present invention and both are incorporated herein by reference.
The parallel plate scintillation detector array comprises a housing with a substantially x-ray transmissive front wall and a plurality of parallel collimator plates disposed orthogonal to the front wall and defining within the housing a plurality of substantially rectangular chambers, each containing a scintillation material. Photoelectrically responsive devices mounted on the top and bottom wall sections of the chamber receive optical radiation produced by excitation of the scintillation material by x-ray or gamma-ray radiation and produce in response thereto electrical output signals representative of the intensity of the impinging radiation.
The present invention represents an improvement in performance and reduction in cost compared to the parallel plate scintillation detector arrays described above.
The scintillation detector array in accordance with the invention comprises a housing having a front wall section substantially transparent to x-ray radiation. A plurality of side walls extending orthogonal to the front wall define adjacent triangular prism shaped chambers having alternate, oppositely disposed bases and containing a scintillation material. A photodetector positioned on the base of each chamber detects photons generated by x-rays which, having passed through the substantially x-ray transparent wall section, excite the scintillation material.
A higher signal-to-noise ratio is one of the performance improvements provided by the above-described inventive array structure. The improvement is due to the elimination of one-half of the total number of light-collecting photodetectors as compared to the parallel plate collimator/channel separator array which typically employs two photodetectors per chamber. More specifically, the higher signal-to-noise ratio results from the halved electrical capacitance associated with the reduced number of photodetectors, typically silicon photodiodes.
Other performance improvements include better compliance of the scintillator body to the chamber walls and greater dimensional uniformity of array chambers. This is achieved by employing accurately machined, wedge-shaped scintillator bars which are force fitted and essentially locked into place individually in each chamber. The resulting structure is a rigid symmetrical array which also exhibits reduced thermal sensitivity generally associated with geometrical distortion of incompletely filled parallel plate arrays. Additionally, the construction of the array provides improved light collection over the parallel plate one-photodiode per chamber array.
The improvements described above are significant since they permit the use of lower x-ray flux while providing superior image quality. This is particularly advantageous in medical tomography applications where it is desirable that the intensity of the x-ray be as small as possible to minimize the exposure of the patient. This is made additionally possible by the increased light-collecting efficacy of the inventive array. The chamber side walls are inclined to the photodetector and are preferably coated with specular material so that light emitted by the scintillator material is focused onto the photodetector.
Significant cost reduction is also provided by the triangular chamber configuration of the array. For example, the number of required photodetectors is reduced by half as is the number of accurately machined slots in the top and bottom array frames into which the collimator plates are fitted to form the triangular prism shaped chambers as compared to the parallel plate array. Cost reduction while retaining high efficacy and structural precision represents a significant advance since in order to obtain high quality imaging the detector array must be precisely constructed to provide a high degree of dimensional uniformity from channel (chamber) to channel.
With respect to precision and uniformity of construction, CT arrays are quite different from the two dimensional x-ray image converting screens which may employ photographic film or a planar array of photodiodes as the recording medium. In such screens, exemplified by U.S. Pat. No. 3,936,645 issued to Iversen on Feb. 3, 1976, the primary concern is to dissect or "discretize" the two-dimensional image conversion screen so as to avoid lateral scatter of x-rays and luminescence which usually limits the contrast and resolution of the image. The precise geometry of the individual cells (serpentine, square, hexagonal, or close-packed spheres, etc.) into which the originally continuous screen is divided is not of any particular significance so long as the intercellular divisions comprise an optically and/or radio-opaque material. In contrast, CT arrays require highly precise scintillator material shapes, scintillator thickness, collimator plate alignments, and reflectivities, just to cite a few examples. Hence the advantage of reducing the number of precision adjustments and fittings required in CT arrays is quite apparent.
These and other advantages and improvements provided by the present invention will be more fully described in the detailed description of the invention.